Sensors using high electron mobility transistors

ABSTRACT

Embodiments of the invention include sensors comprising high electron mobility transistors (HEMTs) with capture reagents on a gate region of the HEMTs. Example sensors include HEMTs with a thin gold layer on the gate region and bound antibodies; a thin gold layer on the gate region and chelating agents; a non-native gate dielectric on the gate region; and nanorods of a non-native dielectric with an immobilized enzyme on the gate region. Embodiments including antibodies or enzymes can have the antibodies or enzymes bound to the Au-gate via a binding group. Other embodiments of the invention are methods of using the sensors for detecting breast cancer, prostate cancer, kidney injury, glucose, metals or pH where a signal is generated by the HEMT when a solution is contacted with the sensor. The solution can be blood, saliva, urine, breath condensate, or any solution suspected of containing any specific analyte for the sensor.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a divisional of U.S. application Ser. No.12/724,117, filed Mar. 15, 2010, which is a continuation-in-part ofInternational Patent Application No. PCT/US2008/076885, filed Sep. 18,2008, which claims the benefit of U.S. Provisional Application Ser. No.60/973,302, filed Sep. 18, 2007, U.S. Provisional Application Ser. No.60/975,907, filed Sep. 28, 2007, and U.S. Provisional Application Ser.No. 60/982,310, filed Oct. 24, 2007, all of which are herebyincorporated by reference herein in their entirety, including anyfigures, tables, or drawings.

BACKGROUND OF THE INVENTION

Chemical sensors can be used to analyze a wide variety of environmentaland bodily gases, aerosols, and fluids for properties of interest. Forexample, exhaled breath condensate (EBC) is widely known to be adiagnostically important bodily fluid that can be safely collected. Inparticular, the breath from deep within the lungs (alveolar gas) is inequilibrium with the blood, and therefore the concentrations ofmolecules present in the breath is highly correlated with those found inthe blood at any given time. Analysis of molecules in exhaled breathcondensate is a promising method that can provide information on themetabolic state of the human body, including certain signs of cancer,respiratory disease, and liver and kidney function. Several differentanalysis methods including gas chromatography (GC), chemiluminescence,selected ion flow tube (SIFT), and mass spectroscopy (MS) have been usedto measure different exhaled biomarkers, including hydrogen peroxide,nitrogen oxide, aldehydes, and ammonia. However, these methods varysignificantly in sensitivity.

Another example of sensing application using body fluid is detectingbreast cancer with saliva. The mortality rate in breast cancer patientscan be reduced by increasing the frequency of screening. Theoverwhelming majority of patients are screened for breast cancer bymammography. This procedure involves a high cost to the patient.Moreover, the use of invasive radiation limits the frequency ofscreening. Recent evidence suggests that salivary testing for markers ofbreast cancer may be used in conjunction with mammography. Saliva baseddiagnostics for the protein c-erbB-2, a prognostic breast marker assayedin tissue biopsies of women diagnosed with malignant tumors, has showntremendous potential. Soluble fragments of the c-erbB-2 oncogene and thecancer antigen 15-3 were found to be significantly higher in the salivaof women who had breast cancer than in those patients with benigntumors. Another recent study concluded that epidermal growth factor(EGF) is a promising marker in saliva for breast cancer detection.

Pilot studies indicate that the saliva test is both sensitive andreliable, and is potentially useful in initial detection and follow-upscreening for breast cancer. However, currently saliva samples aretypically obtained from a patient in a dentist's office then sent to atesting lab; it typically takes a few days to get the test results.

To fully realize the potentials of sensors for environmental, healthrelated, chemical and biomedical applications, technologies are neededthat will enable easy, sensitive, and specific detection of chemical orbiomolecules at home or elsewhere. It is also desirable that a testingdevice allows concomitant wireless data transmission into preprogrammeddestinations, such as transmitting breast cancer testing results to adoctor or clinic. If inexpensive technologies that can detect andwirelessly transmit testing results for environmental, health related,chemical and biomedical applications can be developed, early diagnosisof cancers or disease can significantly lower mortality and the cost ofhealth care. In addition, real-time wireless remote sensing forchemicals in the environment may reduce the incidence of disasters byalerting to a chemical hazard.

BRIEF SUMMARY

High electron mobility transistors (HEMTs), and the particularlyexemplified AlGaN/GaN HEMTs, are a key component of the sensorsaccording to embodiments of the invention. AlGaN/GaN HEMTs withspecified surface functionality perform as sensors to detect variousmolecules, including biomarkers, of interest in bodily fluid samples.

For example, embodiments of the invention are directed to surfacefunctionalized AlGaN/GaN HEMT based sensors that can detect prostatecancer, breast cancer, pH, mercury, copper, glucose, and/or evidence ofacute kidney injury or renal failure in samples of exhaled breathcondensate, saliva, urine, blood, or other fluids. In certainembodiments, the devices according to the invention can wirelesslytransmit results in order to facilitate rapid analysis of the results.

One embodiment of the invention is a device for detecting breast cancerthat includes a gold-gated AlGaN/GaN HEMT functionalized with anantibody to an antigen associated with breast cancer as the capturereagent. A method for detection of breast cancer in a saliva sample isalso exemplified.

A second embodiment of the invention is a device for detection of heavymetals. The device includes a gold-gated AlGaN/GaN HEMT functionalizedwith a chelating agent as the capture reagent. A method for detection ofheavy metals in solution comprises analyzing a sample with a gold-gatedAlGaN/GaN HEMT functionalized with a chelating agent such asthioglycolic acid (HSCH₂COOH), cysteamine (NH₂CH₂CH₂SH),1,2-ethanedithiol (HSCH₂CH₂SH), dimercaprol (BAL),diaminoethanetetraacetic acid (EDTA), 2,3-bis-sulfanylbutanedioic acid(DMSA), or 2,3-dimercapto-1-propanesulfonic acid (DMPS).

A third embodiment of the invention is a device for detecting changes inpH. The device includes an AlGaN/GaN HEMT having a gate dielectriccoating, for example a thin Sc₂O₃ layer, used as the capture reagent.The device can further include a cooling element to obtain exhaledbreath condensates for testing. A method for detecting the pH of exhaledbreath condensate is also disclosed.

A fourth embodiment is a device for detecting prostate cancer using acapture reagent formed via carboxylate succinimdyl ester bound prostatespecific antigen (PSA) antibodies linked to thioglycolic acidimmobilized on a gold-coated gate of an AlGaN/GaN HEMT. A method ofdetecting prostate cancer by analysis of PSA in a sample is alsodisclosed.

A fifth embodiment is a device for detecting acute kidney injury orrenal failure where a gold-gated AlGaN/GaN HEMT functionalized withhighly specific KIM-1 antibodies through a self-assembled monolayer ofthioglycolic acid acts as the detector. A method to detect KIM-1 in asample is also disclosed.

A sixth embodiment is a device for detecting glucose in exhaled breathcondensate. The device includes an AlGaN/GaN HEMT having a nanorod arrayselectively grown on the gate area of the HEMT that immobilizes glucoseoxidase (GOx). The nanorods may be, for example, metal oxide and nitridebased nanorods. A method to detect glucose in exhaled breath condensateis also disclosed.

In accordance with embodiments of the present invention, normalizeddetection is provided.

BRIEF DESCRIPTION OF THE FIGURES

FIGS. 1A-1D show the chemical structures for BAL, EDTA, DMSA, and DMPS,respectively.

FIG. 2 shows a schematic representation of an exemplary embodiment of anAlGaN/GaN HEMT sensor of the present disclosure.

FIG. 3A shows a scanning electron microscope (SEM) image of an exemplarygateless HEMT.

FIG. 3B shows a schematic representation of an exemplary HEMT with anon-native oxide as a gate dielectric layer.

FIG. 4A is a plot showing the effects of EBC exposure, in the form ofmultiple exhaled breaths (each breath <1 sec), on the current change.

FIG. 4B is a plot showing the effects of ventilation strength on theHEMT current.

FIGS. 5A and 5B are photographs of the contact angle of a water drop onthe surface of bare Au (FIG. 5A) and thioglycolic acid functionalized Au(FIG. 5B).

FIG. 6 illustrates a plot of changes in HEMT drain-source current forbare Au-gate and Au-gate with thioglycolic acid functionalizationexposed to 10⁻⁵ M Hg²⁺ ion solutions when using an AlGaN/GaN HEMT sensorof the invention.

FIG. 7A shows time dependent response of the drain current as a functionof Hg²⁺, Cu²⁺, and Pb²⁺ ion concentrations for a bare Au-gate AlGaN/GaNHEMT sensor.

FIG. 7B shows time dependent response of the drain current as a functionof Hg²⁺, Cu²⁺, and Pb²⁺ ion concentrations for a thioglycolic acidfunctionalized Au-gate AlGaN/GaN HEMT sensor.

FIG. 8A shows drain current changes in response to Hg²⁺ and Cu²⁺ ions asa function of the ion concentration for the bare Au-gate AlGaN/GaN HEMTsensor.

FIG. 8B shows drain current changes in response to Hg²⁺ and Cu²⁺ ions asa function of the ion concentration for the thioglycolic acidfunctionalized Au-gate AlGaN/GaN HEMT sensor.

FIG. 9 shows a plan view photograph of a multiple cell AlGaN/GaN HEMTsensor.

FIG. 10 shows time dependent change in the drain current in response toNa⁺ and Mg²⁺ with a bare Au-gated HEMT and a thioglycolic acidfunctionalized Au-gated HEMT sensor of the invention.

FIG. 11A shows recyclability for the bare Au-gate.

FIG. 11B shows recyclability for the thioglycolic acid functionalizedAu-gate surface.

FIG. 12A shows a plan view photomicrograph of a completed HEMT device ofthe invention with a 5-nm Au film in the gate region.

FIG. 12B shows a schematic of an AlGaN/GaN HEMT of the invention, wherethe Au-coated gate area was functionalized with PSA antibody onthioglycolic acid.

FIG. 13A shows I-V characteristics of an AlGaN/GaN HEMT sensor of theinvention before and after PSA incubation.

FIG. 13B shows drain current over time for PSA when sequentially exposedto PBS, BSA, and PSA when using an AlGaN/GaN HEMT sensor of theinvention.

FIG. 14A shows drain current over time for PSA from 10 pg/ml to 1 ng/mlwhen using an AlGaN/GaN HEMT sensor of the invention.

FIG. 14B illustrate change of source and drain current versus differentconcentrations from 10 pg/ml to 1 μg/ml of PSA using an AlGaN/GaN HEMTsensor of the invention.

FIG. 15 shows the drain current over time using a HEMT with a gatedielectric of a non-native oxide and a fixed source-drain bias of 0.25 Vfor pH from 3-10.

FIG. 16 shows the drain current over time using a HEMT with a gatedielectric of a non-native oxide at fixed source-drain bias of 0.25 Vfor pH from 7-8.

FIG. 17A shows a plan view photomicrograph of a completed device with a5-nm Au film on the gate region.

FIG. 17B shows a schematic device cross section where the Au-coated gatearea was functionalized with KIM-1 antibody on thioglycolic acid.

FIG. 18 shows I_(DS)−V_(DS) characteristics of an HEMT in both PBSbuffer and 100 ng/ml KIM-1.

FIG. 19 shows time dependent current signal when exposing the HEMT to 1ng/ml and 10 ng/ml KIM-1 in PBS buffer.

FIG. 20 shows the current change for a HEMT as a function of KIM-1concentration.

FIGS. 21A-21D show field emission SEM images of ZnO nanorod arrays grownon the Si substrate spin-coated with different size precursors(Preparation time of ZnO nanoparticle precursors from (a) to (d): 0.5,1, 1.5, and 2 h, respectively).

FIG. 22A shows a schematic of ZnO nanorod gated AlGaN/GaN HEMT.

FIG. 22B shows a SEM image of ZnO nanorod gated AlGaN/GaN HEMT. Upperright inset shows HRTEM image of a ZnO nanorod array grown on the gatearea with different scales.

FIG. 23 shows plot of drain current versus time with successive exposureof glucose from 500 pM to 125 μM in 10 mM phosphate buffer saline with apH value of 7.4.

FIG. 24 shows plot of change of drain current as a function of glucoseconcentrations from 500 pM to 125 μM in 10 mM phosphate buffer salinewith a pH value of 7.4.

FIG. 25A shows a schematic diagram of a differential amplifier circuitthat can be used to output the signal from the response of a normalizedtemperature sensor in accordance with an embodiment of the presentinvention.

FIG. 25B shows a plan view schematic representation of contact pads fora normalized sensing device in accordance with an embodiment of thepresent invention.

FIG. 26A shows a cross-sectional view of an active and control sensorconfiguration for a normalized sensing device for heavy metal detectionin accordance with one embodiment of the present invention.

FIG. 26B shows a cross-sectional view of an active and control sensorconfiguration for a normalized sensing device for prostate cancerantigen detection in accordance with one embodiment of the presentinvention.

FIG. 26C shows a cross-sectional view of an active and control sensorconfiguration for a normalized sensing device for pH detection inaccordance with one embodiment of the present invention.

FIG. 26D shows a cross-sectional view of an active and control sensorconfiguration for a normalized sensing device for kidney injurymolecule-1 detection in accordance with one embodiment of the presentinvention.

FIG. 26E shows a cross-sectional view of an active and control sensorconfiguration for a normalized sensing device for glucose detection inaccordance with one embodiment of the present invention.

FIG. 26F shows a cross-sectional view of an active and control sensorconfiguration for a normalized sensing device for breast cancer antigendetection in accordance with one embodiment of the present invention.

FIG. 27 shows a plot of diode current vs. bias voltage comparing arelated art differential sensor to a normalized sensor in accordancewith an embodiment of the present invention.

DETAILED DISCLOSURE

One shortcoming of HEMT sensors has been a lack of selectivity todifferent analytes due to the chemical inertness of the HEMT surface.Sensor devices according to embodiments of the invention solve thissensitivity problem by functionalization of the gate surface withcapture reagents.

The sensor devices of the subject invention can be used with a varietyof samples having environmental and/or bodily origins, including saliva,urine, blood, breath (including exhaled breath condensates) and othersamples. For example, in certain embodiments of the invention, mercuryor cancer detection is improved. Additionally, sensors according toembodiments of the invention can be re-used, without substantialdiminishment of efficacy.

Group III-N based wide bandgap semiconductors are used as sensitivechemical sensors, especially when made with piezoelectric materials.GaN/AlGaN high electron mobility transistors (HEMTs) form a high densityelectron sheet carrier concentration channel induced by piezoelectricpolarization of the strained AlGaN layer and spontaneous polarization ofthe different ionic strength between the GaN and AlGaN layer. Theconducting 2-dimensional electron gas (2DEG) channel of GaN/AlGaN basedHEMTs is very close to the surface and extremely sensitive to theambient environment, allowing enhanced detection sensitivity.

GaN-based wide energy bandgap semiconductor material systems areextremely chemically stable; this stability is minimally degraded byadsorbed cells. The bond between Ga and N is ionic and proteins easilyattach to the GaN surface. This is an important factor for preparationof a sensitive biosensor having a useful lifetime.

The HEMT sensors of the subject invention can be used to detect, forexample, gases, ions, pH values, proteins, and/or DNA with goodselectivity by modification of the surface in the gate region of theHEMT. HEMT structures can be used in microwave power amplifiers as wellas gas and liquid sensors because of their high 2DEG mobility andsaturation velocity.

In certain embodiments of the invention, a 2-dimensional electron gas(2DEG) at the interface of AlGaN/GaN heterostructures is formed throughthe hetero junction of AlGaN and GaN, which have different bandgaps. The2DEG channel is connected to an Ohmic-type source and drain contacts.The source-drain current is modulated by a third contact, aSchottky-type gate, on the top of the 2DEG channel. For sensingapplications, the third contact is affected by the sensing environment,i.e. the sensing targets change the charges on the gate region and thebehavior of the gate. When analytes accumulate on the gate area, the netcharge on the HEMT surface is changed. The net surface charge alters the2DEG concentration. This electrical detection technique is simple, fast,and convenient.

The detecting signal from the gate can be amplified through thedrain-source current, making the sensor very sensitive. The electricsignal can be easily quantified, recorded and transmitted, unlikefluorescence detection methods that need human inspection and aredifficult to be precisely quantified and transmitted.

Gateless HEMT structures can distinguish liquids with differentpolarities and can quantitatively measure pH over a broad range. Thesensing mechanism for chemical adsorbates in piezoelectric materialsoriginates from compensation of the polarization induced bound surfacecharge by interaction with the polar molecules in liquids. In gatelessAlGaN/GaN heterostructure transistors, the native oxide on the nitridesurface is responsible for the pH sensitivity of the response toelectrolyte solutions. By coating a thin non-native metal oxide, forexample Sc₂O₃, on the gate sensing area, more sensitive and reproduciblehydronium ion, H₃O⁺ (pH) sensing is achieved. In embodiments of theinvention, the non-native metal oxide on the gate sensing area of theHEMT is referred to as a gate dielectric layer.

Usually, it is difficult to control the compositions and thickness ofnative oxides. For embodiments of the invention, a metal oxide such asSc₂O₃ is grown by a molecule beam epitaxy system with excellentcomposition and thickness control. In certain embodiments of theinvention, other gate dielectric layers, such as metal nitrides, can beused rather than metal oxide dielectric layers. The pH response of anoxide/nitride interface can be modeled in terms of formation of hydroxylgroups that lead to a hydronium ion concentration (pH) dependent netsurface change with a resulting change in voltage drop at thesemiconductor/liquid interface.

According to one embodiment of the subject invention, real-timedetection of the pH of exhaled breath uses a breathing tube and ice bathwith an AlGaN/GaN HEMT. The breathing tube samples exhaled breath andthe ice bath condenses the sample that is applied to the AlGaN/GaN HEMT.

In one embodiment, the device may include an AlGaN/GaN HEMT that isoperably coupled to a thermal electric cooling device, which condensesexhaled breath samples. The thermal water vapor and volatile organiccompounds from the exhaled breath condensate change the surface chargeon the HEMT, thus changing the current flowing in the HEMT device for afixed applied bias voltage. In one embodiment, an exhaled breathcondensate (EBC) biosensor of the present disclosure can be handheld,low in cost, and capable of real-time detection without consumablecarrier gases.

Biologically modified field effect transistors (bioFETs), either atconventional or nano-dimensions, can directly detect biochemicalinteractions in aqueous solutions for a wide variety of biosensingapplications. To enhance the practicality of bioFETs, a device accordingto embodiments of the invention is sensitive to biochemical interactionson its surface that is functionalized to probe specific biochemicalinteractions. In one embodiment, the device is stable in aqueoussolutions across a range of pH and salt concentrations. In otherembodiments, the gate region of the device is covered with capturereagents for molecules of interest. The conductance of the devicechanges as interaction occurs between these capture reagents andappropriate species (the molecules of interest) in a sample.

In one embodiment of the invention, a saliva based breast cancerdetector is functionalized in the gate region with chemicals that canbind (or otherwise interact with) breast cancer markers. The gate regionof the HEMT can be a few nanometers to a few millimeters in size. Anarray of the HEMT sensors can be fabricated on a single chip. Each HEMTcan be functionalized with a capture reagent for a breast cancer marker.A set of testing results can be obtained from a series of differentcapture reagents. Simultaneous breast cancer detections with differentcapture reagents can increase the accuracy of the cancer detection.

Because the surface of AlGaN is extremely inert and difficult tooxidize, a thin layer of gold of, for example, about 5 nm can be used asan intermediate layer between AlGaN and certain capture reagents used inembodiments of the invention. A molecule containing a thiol group can beimmobilized on the Au surface by an Au—S bond. Other functional groupscan then bind with a capture reagent. These other functional groups canbe, for example, alcohol, aldehyde, carboxylic acid, phosphate or aminegroups. The immobilized capture reagents can bind with breast cancerbiomarkers (or other target molecules). In one embodiment an Au-gatedGaN/AlGaN HEMT is used as a sensor for the detection of breast cancermarkers in a saliva sample.

In some embodiments, sensors comprise chemical adsorbates on AlGaN/GaNHEMTs where detection originates from compensating or inducing chargesat the AlGaN/GaN interface due to polar molecules in the liquids bondedto the AlGaN/GaN surface. In certain embodiments, the device isfunctionalized at the AlGaN/GaN/HEMT surface having an Au-coated gateregion by chemicals selected for their interaction with a target beingdetected.

According to one embodiment of the invention, thioglycolic acid can beused to assist in functionalizing an AlGaN/GaN HEMT sensor. For example,a self-assembled monolayer of thioglycolic acid can be adsorbed onto anAu-gate due to interaction between gold and the thiol-group. Followingplacement of the thioglycolic acid on the sensor surface, a specificfunctionality of interest may be conjugated to the surface, where thefunctionality is a capture reagent for a specific target being sensed.

According to certain embodiments of the invention, a sensor can includea synthetic or natural compound as a capture reagent with the ability toassociate with a desired target molecule, such as a biomarker. Thecapture reagent associates with the desired target molecule byinteracting with the target molecule in a way that is detectable by theHEMT. The capture reagent may associate with the target molecule bybinding with the target molecule, but embodiments are not limitedthereto.

The capture reagents of certain embodiments of the invention includenaturally occurring and/or synthetic compounds that preferably displayhigh specificity and sensitivity to a target molecule of interest.Suitable compounds include, but are not limited to, antibodies,proteins, and aptamers that can associate with a biomarker. The term“biomarker” refers to a biochemical in the body with a particularmolecular trait that makes it useful for diagnosing a condition,disorder, or disease, and for measuring or indicating the effects orprogress of a condition, disorder, or disease. Antibodies are proteinmolecules that are typically composed of heavy and light polypeptideamino acid chains held together with disulfide bonds. These highlyspecialized proteins are able to recognize and selectively bind certaintypes of antigen molecules. In embodiments of the invention, a sensoremploys antibodies to detect specific antigens.

In an embodiment of the invention, the chemistry of the system occursalong a conductive layer, for example, a gold layer. The conductivelayer supports the propagation of a high frequency test signal and iscapable of binding to (or otherwise associating with) a target molecule,which is typically an antigen or other analyte. In one embodiment,thioglycolic acid bonds the Au layer to antibodies for breast cancerantigens including (but not limited to) EGF, c-erbB-2 and CA15-3 insaliva, where the thiogylcolic acid forms a self-assembled monolayer onthe gold surfaces. Upon binding of the immobilized antibody to anantigen, the gate potential of the HEMT changes, resulting in a changein current of the HEMT at fixed bias voltage. This change in currentallows identification and, preferably, quantification of the amount ofthe target molecule, for example a cancer biomarker, in the sample.

One embodiment of the invention is a portable or hand-held saliva basedbreast cancer sensor. Other embodiments of the invention are sensors toanalyze other bodily fluids or excretions such as breath, urine orblood. Advantages of the sensors include fast response time for results,portability and low cost. In one embodiment, a chemical sensor array canbe integrated with wireless communication circuits for remote sensorapplications. For example, a digital signal cancer detector canwirelessly send the testing results directly to a user's doctor.

In another embodiment of the invention, evidence of kidney injury isdetected by an HEMT functionalized with thioglycolic acid couplingkidney injury molecule-1 (KIM-1) antibodies to a gold surface of thegate of the HEMT. When in the presence of KIM-1, the gate potential ofthe HEMT changes, resulting in a current change in the HEMT at fixedbias voltage. This change in current can be used to detect and,preferably, quantify KIM-1 biomarker present in a sample.

In yet another embodiment of the invention, the sensor is used to detectheavy metals. Heavy metal detection can involve an HEMT functionalizedwith a densely coated capture reagent. Hg²⁺, Cu⁺² and Pb²⁺ detection canbe achieved according to the invention. One embodiment is a method inwhich chelating agents remove heavy metal ions from a sample, wherechelating ligands and metal ions bind to form metal complexes, normallycalled “chelation.” A strong chelating agent is dimercaprol (BAL), whichcontains two thiol groups capable of reacting with arsenic, lead andmercury. Other widely used chelating agents includediaminoethanetetraacetic acid (EDTA), 2,3-bis-sulfanylbutanedioic acid(DMSA), and 2,3-dimercapto-1-propanesulfonic acid (DMPS). FIGS. 1A-1Dshow the chemical structures for BAL, EDTA, DMSA, and DMPS,respectively.

In one embodiment of the invention, Hg²⁺, Cu⁺² or Pb²⁺ is detected whenchelating agents used as the capture reagent immobilize the metal on theHEMT surface. The surface of AlGaN can have a thin layer (˜5 nm) of goldbetween the AlGaN surface and the chelating agent. Gold permitsdeposition of any chelating agent comprising a thiol group on thesurface through Au—S bonding. The thiol, amine, and carboxyl groups ofthe bound chelating agents bind heavy metal ions to the surface of theHEMT.

FIG. 2 shows a schematic of an exemplary embodiment of an AlGaN/GaN HEMTsensor according to an embodiment of the invention. Thefunctionalization is an Au-coated gate area with thioglycolic acid,HSCH₂COOH, for Hg(II) detection. A self assembled monolayer ofthioglycolic acid molecules is adsorbed onto the gold gate by a S—Aubond between the gold surface and thiol-group. The immobilized carboxylgroups function as capture reagents to capture Hg²⁺, Cu⁺², and/or Pb²⁺ions. Alternative binding groups that can function as capture reagentsare derived, for example, from cysteamine (NH₂CH₂CH₂SH) or1,2-ethanedithiol (HSCH₂CH₂SH).

In one embodiment, the gold-gated region is functionalized withchelating agents immobilized on the HEMT surface, such as BAL, EDTA,DMSA, and DMPS. One portion of the chelating agent binds to the Ausurface and the other portion functions as a capture reagent of heavymetals by chelating with heavy metals, such as Hg²⁺, Cu⁺², or Pb²⁺. Thecharge of the metal ions affects the gate potential of HEMTs. The changein current in the HEMT at fixed bias voltage allows detection and,preferably, quantification of the amount of the heavy metal ions in asample.

In one embodiment, the device is a portable or hand-held traceheavy-metal sensor for environmental and health related applications.The sensor can detect heavy metals in aqueous solution including breathcondensate, urine or blood. Advantages of the sensing device includefast response time, portability and low cost. In one embodiment, a heavymetal detector can be used as a wireless based sensor to transmit adigital signal of the test results directly to a recipient.

Another embodiment of the invention is a pH meter for fluids such asbreath, saliva, urine or blood. Gates of HEMTs can be functionalizedwith noble metal oxides for detecting proton and hydroxide ions. In oneembodiment, a Sc₂O₃ gate dielectric is formed on AlGaN/GaN HEMTs toprovide high sensitivity for detecting changes in pH of electrolytesolutions. HEMTs with Sc₂O₃ exhibit a linear change in current of 37μA/pH between a pH range of 3 to 10. The HEMT pH sensors are stable witha resolution of <0.1 pH over the entire pH range. The HEMTs can be usedto monitor solution pH changes between 7 and 8, a range of interest fortesting human blood.

FIG. 3A shows a scanning electron microscope (SEM) image of a HEMT witha gate dielectric layer. FIG. 3B shows a schematic diagram of a HEMTwith a gate dielectric layer (labeled as oxide). FIG. 4A shows theeffects of EBC exposure, in the form of multiple exhaled breaths (whereeach breath <1 second), on the current. FIG. 4B shows the effects ofventilation strength on the HEMT current, where the duration of thebreath is 5 seconds.

In other embodiments of the invention, a nanorod gated AlGaN/GaN HEMT isa detector for glucose. The nanorod arrays can be selectively grown onthe gate area to immobilize glucose oxidase (GOx). The nanorods can be,for example, metal oxide and/or nitride based nanorods. Nanorod metaloxides include, but are not limited to, SnO, TiO₂, GaN, MgO, ZnMgO, andIn₂O₃ nanorods. For example, one-dimensional ZnO nanorods on the gatearea result in a very high specific surface area with high surface tovolume ratio and provide favorable micro-environments for theimmobilization of GOx.

The AlGaN/GaN HEMT drain-source current has a rapid response, of lessthan 5 seconds, when glucose in a buffer with a pH value of 7.4 is addedto the GOx immobilized ZnO nanorods surface. A wide range of glucoseconcentrations from to 0.5 nM to 125 μM can be detected. For example onesensor according to an embodiment of the invention exhibited a linearrange from 0.5 nM to 14.5 μM with a limit of detection of 0.5 nM.

Following are examples that illustrate embodiments of the invention.These examples should not be construed as limiting. All percentages areby weight and all solvent mixture proportions are by volume unlessotherwise noted.

EXAMPLE 1 Selective Detection of Hg(II) ions from Cu(II) and Pb(II)

Hg²⁺ and Cu²⁺ ions are easily detected with sensors fabricated withAu-gated and thioglycolic acid functionalized Au-gated GaN/AlGaN HEMTs.

The HEMT structures consisted of a 2 μm thick undoped GaN buffer and 250Å thick undoped Al_(0.25)Ga_(0.75)N cap layer. The epi-layers were grownby molecular beam epitaxy system on 2″ sapphire substrates at SVTAssociates. Mesa isolation was performed with an Inductively CoupledPlasma (ICP) etching with Cl₂/Ar based discharges at −90 V dc self-bias,ICP power of 300 W at 2 MHz, and a process pressure of 5 mTorr. Ohmiccontacts of 50×50 μm² separated with gaps of 10, 20, and 50 μm wereformed by e-beam deposition of Ti/Al/Pt/Au patterns by lift-off andannealed at 850° C. for 45 sec under flowing N₂ for source and drainmetal contacts. A 5-nm thin gold film was deposited as the gate metalfor two sets of sample sensors. One sensor had a bare Au-gate and theother sensor had an Au-gate that was functionalized with aself-assembled monolayer of thioglycolic acid. An increase in thehydrophilicity of the surface treated with thioglycolic acidfunctionalization was confirmed by contact angle measurements of a waterdrop of the surface of bare Au (see FIG. 5A) and thioglycolic acidfunctionalized Au (see FIG. 5B), which showed a change in contact anglefrom 58.4° to 16.2° after the surface treatment. A 500-nm-thickpoly(methyl methacrylate) (PMMA) film was used to encapsulate thesource/drain regions, with only the gate region exposed to allow theliquid solutions to access the bare Au-gate or functionalized Au-gatesurface. The source-drain current-voltage characteristics were measuredat 25° C. using an Agilent 4156C parameter analyzer with the Au-gatedregion exposed to different concentrations of Hg²⁺, Cu²⁺, Pb²⁺, Mg²⁺ orNa⁺ solutions. AC measurements were performed to prevent sideelectrochemical reactions with modulated 500-mV bias at 11 Hz.

A schematic cross-section of the device with Hg²⁺ ions bound tothioglycolic acid functionalized on the gold gate region is shown inFIG. 2. A self assembled monolayer of thioglycolic acid molecule wasadsorbed onto the Au-gate due to strong interaction between gold and thethiol-group for the functionalized sensors. Excess thioglycolic acidmolecules were rinsed from the monolayer using DI water. XPS andelectrical measurements confirmed a high surface coverage ofthioglycolic acid molecules with Au—S bonding formation on the AlGaNsurface.

FIG. 6 shows the change in drain current of a bare Au-gated AlGaN/GaNHEMT sensor and a thioglycolic acid functionalized AlGaN/GaN HEMT sensorexposed to 10⁻⁵ M Hg²⁺ ion solutions as compared to being exposed to DIwater (I_(H) ₂ _(O)−I₁₀ ⁻⁵ _(M of HgCl) ₂ ). The drain current of bothsensors decreased after exposure to Hg² ion solutions. The drain currentreduction of the thioglycolic acid functionalized AlGaN/GaN HEMT sensorsexceeded that of the bare Au-gate sensor by almost 80%. Though not to bebound by theory, the mechanisms of the drain current reduction for bareAu-gate and thioglycolic acid functionalized AlGaN/GaN HEMT sensors areprobably quite different. For the thioglycolic acid functionalizedAlGaN/GaN HEMT, the thioglycolic acid molecules on the Au surface alignwith carboxylic acid functional groups extending toward the solution.The carboxylic acid functional group of the adjacent thioglycolic acidmolecules can form chelates (R—COO⁻(Hg²⁺)⁻OOC—R) with the Hg²⁺ ions.Upon chelation, one would expect the charges of trapped Hg²⁺ ion in theR—COO⁻(Hg²⁺)⁻OOC—R to change the polarity of the thioglycolic acidmolecules. Because Hg²⁺ ions were used in the experiments, no Au-mercuryamalgam is expected to form on the bare Au-surface.

FIGS. 7A and 7B show time dependence of the drain current for the twotypes of sensors for detecting Hg²⁺, Cu²⁺, and Pb²⁺ ions. Both type ofsensors showed very short response time (less than 5 seconds), whenexposed to Hg²⁺ ion solution. The limits of detection for Hg²⁺ iondetection for the bare Au-gate and thioglycolic acid functionalizedsensor were 10⁻⁶ and 10⁻⁷ M, respectively. Neither sensor could detectPb²⁺ ions. For the Cu²⁺ ions, the detection limit of the thioglycolicacid functionalized sensor was around 10⁻⁷ M, while the bare Au-gatecould not detect the Cu²⁺ ions as shown in FIG. 7.

FIGS. 8A and 8B show the drain current changes in response to Hg²⁺ andCu²⁺ ions as a function of the ion concentration for the two differentsurfaces. The difference in the response between the bare Au-gate andthe thioglycolic acid functionalized sensor offers the possibility forselective detection for Hg²⁺ and Cu²⁺ ions presented in a singlesolution with a sensor chip containing both types of sensors, as shownin FIG. 9. The dimension of the active area of the AlGaN/GaN HEMT sensoris less than 50 μm×50 μm, and the sensors can be fabricated as an arrayof individual sensors. The fabrication of both sensors is identicalexcept for the thioglycolic acid functionalized sensor, which has anadditional functionalization step. This step can be accomplished with amicro-inkjet system to locally functionalize surfaces. The bare Au-gateand thioglycolic acid functionalized sensors also showed excellentsensing selectivity (over 100 times higher selectivity) over Na⁺ andMg²⁺ ions. As illustrated in FIG. 10, there was almost no detection ofNa⁺ and Mg²⁺ ions for both types of sensors with 0.1 M concentrations.

Most semiconductor based chemical sensors are not reusable. The bareAu-gate and thioglycolic acid functionalized sensors showed very goodreusability, as shown in FIGS. 11A and 11B, respectively. After a simplerinse with DI water, the sensors can be reused for Hg²⁺ ion detectionrepeatedly and the responses to different ionic solutions remainunchanged. The stability of thioglycolic acid functionalized Au surfaceis affected by several factors, such as oxygen level, light, and initialpacking quality. The subject devices were stored in nitrogen ambient andrepeatedly used over a couple of weeks without substantial diminishmentof efficacy.

The Hg²⁺/Ca²⁺ sensor can operate at 0.5 V of drain voltage and 2 mA ofdrain current. However, the operation voltage and device size can befurther reduced to minimize the power consumption to μW range. Thesensor can be integrated with a commercially available hand-heldwireless transmitter to realize a portable, fast response and highsensitivity Hg²⁺ and Cu²⁺ ion detector.

In summary, bared Au-gate and thioglycolic acid functionalized AlGaN/GaNHEMT sensors have demonstrable ability to detect heavy ions. The bareAu-gate sensor was sensitive to Hg²⁺, and thioglycolic acidfunctionalized sensors could detect both Hg²⁺ and Cu²⁺ ions. Byfabricating an array of the sensors on a single chip and selectivelyfunctionalizing some sensors with thioglycolic acid, a multi-functionalspecific detector can be fabricated. Such a sensor array can be used toquantitatively detect Hg²⁺ ions in Cu²⁺ ion solution or Cu²⁺ in Hg²⁺ ionsolution. Both bare Au-gate and thioglycolic acid functionalized sensorscan be repeatedly used after a simple DI water rinse.

EXAMPLE 2 Detection of Prostate Specific Antigen

Functionalized of Au-gated AlGaN/GaN HEMTs of the invention were used todetect prostate specific antigen (PSA). The PSA was specificallyrecognized through PSA antibody, anchored to the gate area in the formof carboxylate succinimidyl ester. A wide range of concentrations fromto 1 μg/ml to 10 pg/ml of PSA was investigated, which is lower than thecut-off value of 2.5 ng/ml that is used as an indication for the need ofbiopsy.

The HEMT structures consisted of a 3 μm thick undoped GaN buffer, a 30 Åthick Al_(0.3)Ga_(0.7)N spacer, and a 220 Å thick Si-dopedAl_(0.3)Ga_(0.7)N cap layer. Epi-layers were grown by rf plasma-assistedMolecular Beam Epitaxy on the thick GaN buffers produced on sapphiresubstrates by metal organic chemical vapor deposition (MOCVD). Mesaisolation was performed with an Inductively Coupled Plasma (ICP) etchingwith Cl₂/Ar based discharges at −90 V dc self-bias, ICP power of 300 Wat 2 MHz, and a process pressure of 5 mTorr. 10×50 μm² Ohmic contactsseparated with gaps of 5 μm were formed by e-beam deposited Ti/Al/Pt/Aupatterned by lift-off and annealed at 850° C. for 45 sec under flowingN₂. Poly(methyl methacrylate) (PMMA) was used to form 400-nm-thick layerencapsulating the source/drain regions, with only the gate regionexposed to allow the liquid solutions to contact the gate surface. Thesource-drain current-voltage characteristics were measured at 25° C.using an Agilent 4156C parameter analyzer with the gate region exposedto solution. AC measurements were performed with modulated 500-mV biasat 11 Hz to prevent side electrochemical reactions.

A plan view photomicrograph of a completed device is shown in FIG. 12Aand a schematic cross-section of the device is shown in FIG. 12B. The Ausurface was functionalized with a specific bifunctional molecule. Here,thioglycolic acid, HSCH₂COOH, was attached to the Au surface in the gatearea as a self assembled monolayer adsorbed on the gold gate.

The devices were first placed in the ozone/UV chamber for 3 minutes andthen submerged in a 1 mM aqueous solution of thioglycolic acid for 24hours at room temperature, resulting in binding of the thioglycolic acidto the Au surface in the gate area with the COOH groups available forfurther chemical functionalization. XPS and electrical measurements weretaken to confirm a high surface coverage and Au—S bonding formation onthe surface. The device was freshly cleaned with deionized water toremove unlinked thioglycolic acids. The carboxylic acid functionalgroups were activated by submerging the device in a 0.1 mM solution ofN,N′-dicyclohexylcarbodiimide (DCC) in dry acetonitrile for 30 minutesand then in a 0.1 mM solution of N-hydroxysuccinimide in dryacetonitrile for 1 hour. These functionalization steps resulted in theformation of succinimidyl ester groups on the gate area of AlGaN/GaNHEMT, as shown in FIG. 12B. The device was incubated in a phosphatebuffered saline (PBS) solution of anti-PSA monoclonal antibody for 18hours before real time measurement of PSA.

After incubation in a PBS buffered solution containing PSA at aconcentration of 1 μg/ml, the device surface was thoroughly rinsed withdeionized water and dried with a nitrogen stream. The electricalproperties of the devices, source and drain current, were measuredbefore and after PSA incubation as shown in FIG. 13A. As previouslydescribed, the electrons in 2DEG channel of the AlGaN/GaN HEMT areinduced by piezoelectric and spontaneous polarization effects. Positivecounter-charges at the AlGaN surface layer are induced by the 2DEG. Anyslight changes in the ambient environment of the AlGaN/GaN HEMT affectthe surface charges of the AlGaN/GaN HEMT. These changes in the surfacecharge are transduced into a change in the concentration of the 2DEG inthe AlGaN/GaN HEMTs, leading to the slight decrease in the conductancefor the device after PSA incubation.

FIG. 13B shows the real time PSA detection in PBS buffer solution usingchange in the source-to-drain current with a constant bias of 500 mV. Nocurrent change can be seen with the addition of buffer solution around100 sec and the addition of nonspecific bovine serum albumin (BSA)around 200 sec, showing relatively high stability of the device andchemical surface modification. In clear contrast, the current changeshowed a rapid response of less than 5 seconds when 10 ng/ml PSA wasintroduced to the antibody on the surface. The abrupt current change,mainly due to the exposure of PSA in a buffer solution, stabilized afterthe PSA thoroughly diffused into the buffer solution.

Further real-time testing for the detection limit of PSA of less diluteconcentrations was carried out as shown in FIG. 14A. Three differentconcentrations of the exposed target PSA in a buffer solution wereobserved from 10 pg/ml to 1 ng/ml. The amplitude of current change forthe device exposed to PSA in a buffer solution was about 3%, asillustrated in FIG. 14B. The clear current decrease of 64 nA as 10 pg/mlof PSA also indicated that the detection limit could be lowered up toseveral pg/ml, showing the promise of a portable electronic biologicalsensor for PSA screening.

As demonstrated herein, through a chemical modification sequence, theAu-gated region of an AlGaN/GaN HEMT structure can be functionalized forthe detection of PSA with a sensitivity of 10 pg/ml in a buffer atclinical concentration. This electronic detection of biomolecules canoccur with a compact sensor chip, which can be integrated with acommercially available hand-held wireless transmitter to realize aportable, fast and high-sensitivity prostate cancer detector.

EXAMPLE 3 Detection of Changes in pH in Electrolyte Solutions

A Sc₂O₃ gate dielectric on AlGaN/GaN HEMTs is shown to provide highsensitivity for detecting changes in pH of electrolyte solutions, and issuperior to the use of native oxide in the gate region.

The HEMT structures consisted of a 2 μm thick undoped GaN buffer and 250Å thick undoped Al_(0.25)Ga_(0.75)N cap layer. The epi-layers were grownby Metal-Organic Chemical Vapor Deposition on 100 mm (111) Si substratesat Nitronex Corporation. The sheet carrier concentration was ˜1×10¹³cm⁻² with a mobility of 980 cm²/V-s at room temperature. Mesa isolationwas achieved by using an ICP system with Ar/Cl₂ based discharges. Ohmiccontacts of 50×50 μm² separated with gaps of 10, 20, and 50 μm wereformed by lift-off of e-beam deposited Ti(200 Å)/Al(800 Å)/Pt(400Å)/Au(800 Å). The contacts were annealed at 850° C. for 45 sec under aflowing N₂ ambient in a Heatpulse 610T system. A 100 Å Sc₂O₃ layer wasdeposited as a gate dielectric through a contact window of SiN_(x)layer. Before oxide deposition, the wafer was exposed to ozone for 25minutes, and heated in-situ at 300° C. for 10 minutes inside the growthchamber. The Sc₂O₃ was deposited by rf plasma-activated MBE at 100° C.using elemental Sc evaporated from a standard effusion cell at 1130° C.and O₂ derived from an Oxford RF plasma source.

For comparison, devices with only the native oxide present in the gateregion and devices with the UV ozone-induced oxide were fabricated. FIG.3A shows a scanning electron microscopy (SEM) image and FIG. 3B shows across-sectional schematic of the completed device. The gate dimension ofthe device is 2×150 μm². The pH solution was applied using a syringeautopipette (2-20 μl).

Prior to the pH measurements, pH 4, 7, and 10 buffer solutions fromFisher Scientific were used to calibrate the electrode and themeasurements at 25° C. were carried out in the dark using an Agilent4156C parameter analyzer to avoid parasitic effects. The pH solution wasmade by the titration method using HNO₃, NaOH, and distilled water. Theelectrode was a conventional Acumet standard Ag/AgCl electrode.

The adsorption of aqueous solution of different pH on the surface of theHEMT affected the surface potential and device characteristics. FIG. 15shows the current at a bias of 0.25V as a function of time from theHEMTs with Sc₂O₃ in the gate region exposed for 150 sec to a series ofsolutions whose pH was varied from 3-10. The current significantlyincreased as the pH was decreased upon exposure to these aqueoussolutions. The change in current was 37 μA/pH. The HEMTs show stableoperation with a resolution of ˜0.1 pH over the entire pH range,illustrating the remarkable sensitivity of the HEMT to relatively smallchanges in concentration of the hydronium ion in solution. Bycomparison, devices with the native oxide in the gate region showed ahigher sensitivity of ˜70 pA/pA but a poor resolution of ˜0.4 pH andshowed delays in response of 10-15 seconds. The delays may result fromdeep traps at the interface between the semiconductor and native oxide,whose density is much higher than at the Sc₂O₃-nitride interface. Thedevices with UV-ozone oxide in the gate region did not show theseincubation times for detection of pH changes and showed similarsensitivities of gate source current as the Sc₂O₃ gate devices (˜40pA/pH), but displayed poorer resolution (˜0.25 pH). FIG. 15 shows thatthe HEMT sensor with Sc₂O₃ gate dielectric is sensitive to theconcentration of the polar liquid and therefore could be used todifferentiate between liquids into which a small amount of leakage ofanother substance has occurred.

The pH range of interest for human blood is 7-8. FIG. 16 shows thechange in current of the HEMTs with Sc₂O₃ at a bias of 0.25V fordifferent pH values in this range. The resolution of the measurement is<0.1 pH. As previously described, the electrons in the 2DEG channel ofthe AlGaN/GaN HEMT are induced by piezoelectric and spontaneouspolarization effects. Positive counter charges at the AlGaN surfacelayer are induced by the 2DEG. Any change in the ambient environment ofthe AlGaN/GaN HEMT affects the surface charges of the device. Thesechanges in the surface charge are transduced into a change in theconcentration of the 2DEG. Different pHs exhibit different degrees ofinteraction with the Sc₂O₃ surface. These results show that using ahigher quality oxide is useful in improving pH resolution.

EXAMPLE 4 Detection of Kidney Injury Molecule-1

The HEMT structure consisted of a 2 μm thick undoped GaN buffer and 250Å thick undoped Al_(0.25)Ga_(0.75)N cap layer. The epi-layers were grownby metal-organic chemical vapor deposition on 100 mm (111) Sisubstrates. Mesa isolation was performed with ICP etching with Cl₂/Arbased discharges at −90 V dc self-bias, ICP power of 300 W at 2 MHz, anda process pressure of 5 mTorr. Ohmic contacts of 50×50 μm² separatedwith gaps of 20 μm were formed by e-beam deposited Ti/Al/Pt/Au patternedby lift-off and annealed at 850° C. for 45 sec under a N₂ stream. A 5-nmthin gold film was deposited as a gate metal and functionalized as aself-assembled monolayer of thioglycolic acid. Poly(methyl methacrylate)(PMMA) was used to form a 500-nm-thick encapsulate of the source/drainregions, with the gate region exposed using e-beam lithography. A planview photomicrograph of a completed device is shown in FIG. 17A.

Before depositing the thioglycolic acid coating, the sample was exposedto UV ozone for 5 minutes to remove surface contamination. Aself-assembled monolayer of thioglycolic acid molecule was adsorbed ontothe Au-gate. Excess thioglycolic acid was rinsed off with PBS buffer. Anincrease in the hydrophilicity of the treated thioglycolic acidfunctionalization surface was confirmed by contact angle measurements,which showed a change in contact angle from 58.4° to 16.2° after thesurface treatment (see e.g., FIGS. 5A and 5B).

The thioglycolic acid surface was treated with monoclonal anti-ratkidney injury molecule-1 (KIM-1) antibody in a solution of 10 mMphosphate buffer containing 4 mM sodium cyano-borohydride with pH 8.8 atroom temperature for 2 hours. This antibody immobilization is based on astrong reaction between the carboxyl group on thioglycolic acid and theamine group on KIM-1 antibody. Excess KIM-1 antibodies were washed fromthe surface using a PBS buffer and the unreacted surface carboxyl groupswere passivated by a blocking solution of 100 mM ethanolamine in 10 mMphosphate buffer with pH 8.8. FIG. 17B shows a schematic devicecross-section with thioglycolic acid followed by KIM-1 antibody coating.The source-drain current-voltage characteristics were measured at 25° C.using an Agilent 4156C parameter analyzer with the KIM-1 antibodyfunctionalized Au-gated region exposed to different concentrations ofKIM-1/PBS buffer. AC measurements were performed to prevent sideelectrochemical reactions with modulated 500-mV bias at 11 Hz.

The source-drain current (I_(DS)) vs. voltage (V_(DS)) for the devices,were measured in PBS buffer and 100 ng/ml KIM-1 in PBS buffer, as shownin FIG. 18. There is a clear conductance decrease with KIM-1 exposureand this suggests that through the selective binding of KIM-1 withantibody, there are charges accumulated at the surface of the HEMT andthese surface charges are transduced into a change in the carrierconcentration of AlGaN/GaN 2DEG, leading to the obvious decrease in theconductance of the device after KIM-1 exposure.

FIG. 19 shows the time dependent source-drain current signal with aconstant bias of 500 mV for KIM-1 detection in PBS buffer solution. Nocurrent change can be seen with the addition of buffer solution around50 sec. This stability excludes the possibility of noise due to themechanical change of the buffer solution. By sharp contrast, the currentchange showed a rapid response in less than 20 seconds when 1 ng/mlKIM-1 was switched to the surface at 150 sec. The abrupt current changedue to the exposure of KIM-1 in a buffer solution stabilized after theKIM-1 thoroughly mixed with the buffer. A 10 ng/ml KIM-1 solution wasthen applied at 350 sec, which was accompanied by a larger signal due tothe higher KIM-1 concentration.

Additional real-time tests were carried out to explore the limits ofdetection of KIM-1 antibody. Referring to FIG. 20, the device wasexposed to 10 pg/ml, 100 pg/ml, 1 ng/ml, 10 ng/ml, and 100 ng/mlindividually with each concentration repeated five times to determinethe standard deviation of source-drain current response for eachconcentration. The limit of detection of this device was 1 ng/ml KIM-1in PBS buffer solution and the source-drain current change isnonlinearly proportional to the KIM-1 concentration. Between each test,the device was rinsed with a wash buffer of 10 μM phosphate buffersolution containing 10 μM KCl with pH 6 to strip the antibody from theantigen. These results suggest that the HEMTs are compatible with AKIbiomarker, KIM-1; are very sensitive relative to other currentlyavailable nano-devices; and are useful for preclinical and clinicalapplications. Similar surface modifications can be applied for detectingother important disease biomarkers and a compact disease diagnosis arraycan be realized for multiplex disease analysis.

EXAMPLE 5 Glucose Detection

The HEMT structure consisting of a 3 μm thick undoped GaN buffer, 30 Åthick Al_(0.3)Ga_(0.7)N spacer, 220 Å thick Si-doped Al_(0.3)Ga_(0.7)Ncap layer was provided. Epi-layers were grown by both molecular beamepitaxy and metal organic chemical vapor deposition (MOCVD) on thick GaNbuffers on sapphire substrates. Mesa isolation was performed with an ICPetching with Cl₂/Ar based discharges at −90 V dc self-bias, ICP power of300 W at 2 MHz, and a process pressure of 5 mTorr. 50×50 μm² Ohmiccontacts separated with gaps of 10 μm were formed by e-beam depositedTi/Al/Pt/Au patterned by lift-off and annealed at 850° C. for 45 secunder flowing N₂.

ZnO nanorods were grown in a solution of 20 mM zinc acetate hexahydrate(Zn(NO₃)₂.6H₂O) and 20 mM hexamethylenetriamine (C₆H₁₂N₄) in a flaskwith polypropylene autoclavable cap at a controlled temperature and pHon the HEMT substrate. Subsequently, the substrate was removed fromsolution, thoroughly rinsed with acetone followed by deionized water toremove any residual salts and dried in air at room temperature. The ZnOnanoparticle size was highly dependent on the nanocrystal seedpreparation time. The diameters of ZnO nanorods can be carefullycontrolled from tens of nanometers to several hundred micrometers by theseed size used to grow the nanorods. FIGS. 21A-21D show the effect ofnanocrystal seed size on the diameters of ZnO nanorods grown on a Si(111) substrate.

By incorporating the nanorods on the HEMT gate sensing area, the totalsensing area increases significantly as shown in FIG. 22A. Theconventional AlGaN/GaN HEMT detects the ambient changes through the“gate sensing area”. This area is defined as gate length×gate width inthe regular HEMT. Although, the gate width can be increased in order togain higher drain current from the transistor, the sensor detectionsensitivity will be the same for the HEMT having both short and longergate width. This is due to the signal and background currentproportionally increasing at the same time. The other dimension of thegate is the gate length. However, increasing the gate length increasesthe parasitic resistance of the HEMT and the drain current decreases.Thus, the detection sensitivity goes down. Therefore, the only way toincrease the sensitivity with the same “gate dimension” is to grow 3Dstructures on the gate sensing area to increase the total sensing areawith the area expansion to the third dimension.

FIG. 22B shows SEM pictures of ZnO nanorods grown on the AlGaN/GaN HEMTgate sensing area. The upper right inset in FIG. 22B shows a closer viewof ZnO nanorod arrays grown on the gate area with different scales. TheZnO nanorods matrix provides a microenvironment for immobilizingnegatively charged GOx while retaining its bioactivity, and passescharges produce during the GOx and glucose interaction to the AlGaN/GaNHEMT.

A GOx solution was prepared with a concentration of 10 mg/mL in the 10mM phosphate buffer saline (pH value of 7.4., Sigma Aldrich). Afterfabricating the device, 5 μl GOx (˜100 units/mg, Sigma Aldrich) solutionwas dropped on the surface of HEMT device. The HEMT device was kept at4° C. in the solution for 48 hours for GOx immobilization on the ZnOnanorod arrays followed by an extensive washing to remove theunimmobilized GOx. The HEMT device was kept in the incubator for 30minutes to make the enzyme active around 37° C.

The target glucose was applied on the device through a syringeautopipette (2-20 μl). The current-voltage characteristics were measuredusing an Agilent 4156C parameter analyzer with the gate region exposed.

FIG. 23 shows the real time glucose detection in PBS buffer solutionusing the drain current change with constant bias of 250 mV. No currentchange can be seen with the addition of buffer solution at around 200sec, showing the specificity and stability of the device. By sharpcontrast, the current change showed a rapid response in less than 5seconds when target glucose was added to the surface. The response isvery linear from 0.5 nM to 14.5 μM and showed an experimental limit ofdetection of 0.5 nM.

Immobilization of the negatively charged GOx on the positively chargedZnO nanorod arrays is maximized around the pH value of 7.4 and reducesto about 80% for pH=5 to 6. Once the pH value is larger than 8, theactivity drops significantly. The human pH value can vary depending onthe health condition, e.g. the pH value for patients with acute asthmawas reported as low as 5.23±0.21 (n=22) as compared to 7.65±0.20 (n=19)for the control subjects. In order to get accurate measurements ofglucose concentration in the EBC, one needs to know the correlation ofthe pH value of the EBC and the sensitivity of GOx functionalized forspecific sensors. The low detection limit of the sensor permitteddilution of <0.1 μ-liter of exhaled breath condensate (EBC) in 100-200μ-liter PBS and the direct measurement of the glucose concentration toeliminate the effect of pH variation of the EBC. FIG. 24 shows thechanges of drain current as a function of glucose concentration. A verygood linear relationship of glucose concentration vs. drain currentchanges was obtained. Because of the fast response time and low volumeof the EBC required, a handheld real-time glucose sensor can be made.

Embodiments are provided for improved temperature response that utilizea normalized diode or field effect transistor (FET) configuration.According to an embodiment, a control sensor and an active sensor arearranged in a common ground configuration. A differential amplifier canbe connected to the normalized HEMT-based sensor to provide an amplifiedoutput of the sensor's response to the target molecule in solution(i.e., the element being detected in the environment). The differentialamplifier amplifies the difference between the two sensors and rejectsthe signal that is common to the inputs. For example, FIG. 25A shows aschematic representation of a normalized sensor output circuit 5 inaccordance with one embodiment of the present invention. The controlsensor 2 is connected to one input of the output circuit 5 and theactive sensor 3 is connected to another input of the output circuit 5.The control sensor 2 and the active sensor 3 are also connected to acommon node providing a common voltage V_(common). In other embodiments,the schematic shown in FIG. 25A can be replaced with any suitabledifferential amplifier circuit. Though not shown in the figures, aninitialization circuit can be included to reset the sensor and/or tobias the output circuit 5. Any suitable biasing circuit can be used.FIG. 25B shows contact pads for the sensor 10 according to oneembodiment. A first contact pad 11 can connect the control sensor to thedifferential amplifier circuit; a second contact pad 12 can connect thecommon node of the control sensor and the active sensor to a groundsignal; and a third contact pad 13 can connect the active sensor to thedifferential amplifier circuit.

In accordance with embodiments of the present invention, both thecontrol and the active sensor are exposed to the ambient temperature.The control (or reference) sensor of an embodiment of the presentinvention has the exact same gate functionalization to semiconductorinterface as the active sensor of the subject device. Further, the gatefunctionalization of the control sensor is covered with another layer, alayer with a similar binding functional group but different terminatedfunctional group, or antibodies not sensitive to the designateddetection. An example where the gate functionalization of the controlsensor is covered with a layer with similar binding functional group butdifferent terminated function group is using methyl mercaptans (HS—CH₃or HS—(CH₂)_(n)—CH₃, where n=0, 1, 2, . . . ) for the reference sensorfor Hg ion detection where thioglycolic acid (HS—CH₂—CH₂—COOH) is usedfor the active sensor.

The additional layer can be a protective layer of metal, dielectric,antibody, Bovine serum albumin (BSA), protein, aptamer, or polymer,which is inert to the element being detected in the surroundingenvironment or inhibits the gate functionalization material of thecontrol sensor from being exposed to the element being detected in thesurrounding environment.

According to an embodiment of the present invention, the sensingresponse signal is output from the potential difference between thecontrol sensor and the active sensor. The source regions of the sensorsare grounded together for diode mode sensing and the drain regions ofthe sensor are floated. If the FET mode is used for the sensing, thedrain current or threshold voltage of the HEMT is used to monitor theconcentration of the element being detected instead of diode currentused in the diode mode sensing.

The normalized configuration provides a built-in control diode to reducefalse alarms due to temperature swings or voltage transients. Since boththe control and the active sensor have the same interface between thesemiconductor material of the HEMT and the gate functionalization, thediode or FET characteristics will be substantially the same regardlessof ambient temperature. Thus, the differences in diode or FETcharacteristics for the two sensors (control and active) occur only intheir exposure to the environment. Specifically, the active sensor willrespond to elements being detected in the environment and the controlsensor will not.

The HEMT amplifies the signal detected from the gate of the activesensor, thereby enabling extremely sensitive sensing. The amplifiedsignal of the active sensor can then be compared to the control sensorsignal through, for example, the differential amplifier circuit of FIG.25A to provide a normalized signal. In addition, embodiments of thepresent invention can accomplish these reductions in false alarms at awide range of temperatures. In certain embodiments, the subject sensorcan reduce false alarms for temperatures between −40° C. and 80° C.

According to embodiments of the present invention, a normalized diodeconfiguration is provided where the control device includes the sameHEMT interface structure as the gate functionalization of the activedevice, but further includes the metal of a final metal layer,dielectrics, antibody, BSA, protein, aptamer, or polymers.

In one embodiment example, referring to FIG. 26A, for heavy metaldetection, the active member 20 of the normalized pair has a gold-basedgate contact 21 with a chelating agent 22 bonded thereto, and thecontrol member 23 of the normalized pair has a gold-based gate contact21 and a protective layer 24. The protective layer 24 can be a metallayer, a dielectric layer, antibody, BSA, protein, aptamer, methylmercaptans (HS—CH₃ or HS—(CH₂)_(n)—CH₃), or a polymer layer. When ametal layer is used for the protective layer 24, it can be the metallayer(s) used for final metal contacts 25 of the device. In certainimplementations of a heavy metal normalized sensor, the control sensorcan be functionalized similarly to the active sensor by coating thereference sensor with the methyl mercaptans (HS—CH₃ orHS—(CH₂)_(n)—CH₃), metal, dielectric, or polymer instead of thioglycolicacid (HS—CH₂—CH₂—COOH).

In another embodiment example, referring to FIG. 26B, for prostatespecific antigen detection, the active member 26 of the normalized pairhas a gold-based gate contact 27 with PSA antibody 28 anchored to thegate area with a binding agent 29 to the gold-based gate contact 27, andthe control member 30 of the normalized pair has a gold-based gatecontact 27 and a protective layer 31. The protective layer 31 can be ametal layer, a dielectric layer, antibodies not sensitive to theprostate specific antigen, BSA, protein, aptamer, methyl mercaptans(HS—CH₃ or HS—(CH₂)_(n)—CH₃), or a polymer layer. When a metal layer isused for the protective layer 31, it can be the metal layer(s) used forfinal metal contacts 32 of the device. In certain implementations of aPSA normalized sensor, the control sensor can be functionalizedsimilarly to the active sensor by coating the gate area of the referencesensor with the metal, dielectric, antibodies not sensitive to theprostate specific antigen, BSA, protein, aptamer, methyl mercaptans(HS—CH₃ or HS—(CH₂)_(n)—CH₃), or polymer.

In yet another embodiment example, referring to FIG. 26C, for detectionof changes in pH in electrolyte solutions, the active member 33 of thenormalized pair has a non-native oxide 34 on the gate region, and thecontrol member 35 of the normalized pair has the non-native oxide 34 anda protective layer 36. The protective layer 36 can be a metal layer,another dielectric layer, or a polymer layer. When a metal layer is usedfor the protective layer 36, it can be the metal layer(s) used for finalmetal contacts 37 of the device.

In another embodiment example, referring to FIG. 26D, for kidney injurymolecule-1 detection, the active member 38 of the normalized pair has agold-based gate contact 39 with KIM-1 antibody 40 anchored to the gatearea with a binding agent 41 to the gold-based gate contact 39, and thecontrol member 42 of the normalized pair has the gold-based gate contact39 and a protective layer 43. The protective layer 43 can be a metallayer, a dielectric layer, antibodies not sensitive to the kidney injurymolecule-1, BSA, protein, aptamer, methyl mercaptans (HS—CH₃ orHS—(CH₂)_(n)—CH₃), or a polymer layer. When a metal layer is used forthe protective layer 43, it can be the metal layer(s) used for finalmetal contacts 44 of the device. In certain implementations of a KIM-1normalized sensor, the control sensor can be functionalized similarly tothe active sensor by coating the gate area of the reference sensor withthe metal, dielectric, antibodies not sensitive to the kidney injurymolecule-1, BSA, protein, aptamer, methyl mercaptans (HS—CH₃ orHS—(CH₂)_(n)—CH₃), or polymer.

In one embodiment example, referring to FIG. 26E, for glucose detection,the active member 45 of the normalized pair has ZnO nanorods 46 on thegate region, and the control member 47 of the normalized pair has theZnO nanorods 46 and a protective layer 48. The protective layer 48 canbe a metal layer, an enzyme layer not sensitive to the glucose, adielectric layer, or a polymer layer. When a metal layer is used for theprotective layer 48, it can be the metal layer(s) used for final metalcontacts 49 of the device.

In yet another embodiment example, referring to FIG. 26F, for breastcancer specific antigen detection, the active member 50 of thenormalized pair has a gold-based gate contact 51 with breast cancerspecific antibody 52, such as EGF, c-erbB-2, and CA15-3, anchored to thegate area with a binding agent 53 to the gold-based gate contact 51, andthe control member 54 of the normalized pair has a gold-based gatecontact 51 and a protective layer 55. The protective layer 55 can be ametal layer, a dielectric layer, antibodies not sensitive to the breastcancer specific antigen, BSA, protein, aptamer, methyl mercaptans(HS—CH₃ or HS—(CH₂)_(n)—CH₃), or a polymer layer. When a metal layer isused for the protective layer 55, it can be the metal layer(s) used forfinal metal contacts 56 of the device. In certain implementations of thenormalized sensor, the control sensor can be functionalized similarly tothe active sensor by coating the gate area of the reference sensor withthe metal, dielectric, antibodies not sensitive to the breast cancerspecific antigen, BSA, protein, aptamer, methyl mercaptans (HS—CH₃ orHS—(CH₂)_(n)—CH₃), or polymer.

Advantageously, by providing both the control device and the activedevice with a same or similar gate functionalization, Schottkycharacteristics exist for both devices. In addition, the inclusion ofthe same gate functionalization to HEMT semiconductor interface for boththe active device and the control device brings the work function of thecontrol device in line with the active device, thereby making theresponse to different temperature ambients in line with each other andreducing the effects of having different responses to the differenttemperature ambients. This effect can be seen in FIG. 27. FIG. 27 showsa plot of the difference of diode current vs. bias voltage between areference (control) and an active sensor for active and control sensorsusing different Schottky metallization at three temperature ambientscompared to an example embodiment where the active and control sensorsuse a same Schottky metallization. It should be noted that these plotsrepresent the current in an ambient environment having no element beingdetected. Therefore, for the sensing device of embodiments of thepresent invention, the temperature and bias dependence of the responsesignal can be minimized and the signal be normalized to indicate onlythe presence of the element being detected (as opposed to temperaturedependence).

Specifically, as shown in FIG. 27, by not providing both the controldevice and the active device with a same or similar gatefunctionalization, the device where the control and active sensors usedifferent Schottky metallization as the gate contact shows sensitivityto temperature changes and applied bias. In contrast, an embodiment ofthe present invention using same Schottky metallization as the gatecontact is capable of maintaining a constant current over change in biasand temperature. As shown in the plot, even where a control diode isused (but does not have the same gate metal as the active diode), thedifference of the diode current from the control and active diode is notzero (normalized). This is due to the gate metal/semiconductor interfaceof the control sensor being different from the active diode. The workfunction of the different metal/semiconductor interfaces is differentand the diode current of the two sensors (control and active) would bedifferent at different bias as well as different temperature.

All patents, patent applications, provisional applications, andpublications referred to or cited herein are incorporated by referencein their entirety, including all Figures and tables, to the extent theyare not inconsistent with the explicit teachings of this specification.

It should be understood that the examples and embodiments describedherein are for illustrative purposes only and that various modificationsor changes in light thereof will be suggested to persons skilled in theart and are to be included within the spirit and purview of thisapplication.

What is claimed is:
 1. A high electron mobility transistor (HEMT)-basedsensor for one or more target molecules in a sample, wherein the HEMT ofthe sensor comprises: a gate region having at least one capture reagentthereon; and a gate dielectric layer comprising nanorods selectivelyformed on the gate region.
 2. The HEMT-based sensor according to claim1, wherein the HEMT is an AlGaN/GaN HEMT.
 3. The HEMT-based sensoraccording to claim 2, wherein the nanorods comprise a metal oxide. 4.The HEMT-based sensor according to claim 3, wherein the metal oxidecomprises at least one of ZnO, SnO, TiO₂, MgO, ZnMgO, Al₂O₃, and In₂O₃.5. The HEMT-based sensor according to claim 4 wherein the capturereagent is immobilized on the nanorods.
 6. The HEMT-based sensoraccording to claim 2, wherein the nanorods are ZnO nanorods.
 7. TheHEMT-based sensor according to claim 6, wherein the capture reagent isimmobilized on the nanorods.
 8. The HEMT-based sensor according to claim1, wherein the nanorods are ZnO nanorods.
 9. The HEMT-based sensoraccording to claim 1, wherein the HEMT is configured to detect the oneor more target molecules in at least one of exhaled breath condensate,saliva, urine, blood, other biological fluids, and other aqueoussolutions.
 10. The HEMT-based sensor according to claim 1, wherein theHEMT comprises a two layer structure of group III-IV semiconductormaterials, the first layer of the two layer structure being a layerhaving a first ionic strength and the second layer of the two layerstructure being a strained layer on the first layer and having a secondionic strength different than the first ionic strength, wherein a highdensity electron sheet carrier concentration channel of a 2-dimensionalgas channel is induced by piezoelectric polarization of the strainedlayer and spontaneous polarization of the different ionic strengthsbetween the first layer and the strained layer, and wherein the sensoris configured such that an output current indicating an amount of targetmolecules is sensed from a drain of the HEMT with a fixed bias voltageat a source of the HEMT, in response to a change of concentration of the2-dimensional gas channel resulting from a change of surface charge onthe gate region of the HEMT when the gate region is exposed to the oneor more target molecules.